Device and method for creating aerosols for drug delivery

ABSTRACT

A drug delivery device and method is disclosed which produces aerosolized particles of pharmaceutically active drug for delivery to a patient by inhalation. The device is comprised of a liquid feeding source such as a channel to which formulation is added at one end and expelled through an exit opening. The feeding channel is surrounded by a pressurized chamber into which gas is fed and out of which gas is expelled from an opening. The opening from which the gas is expelled is positioned directly in front of the flow path of liquid expelled from the feeding channel. Various parameters are adjusted so that pressurized gas surrounds liquid flowing out of the feeding channel in a manner so as to maintain a stable capillary microjet of liquid until the liquid exits the pressure chamber opening and is aerosolized. The aerosolized particles having a uniform diameter in the range of about 1 to 5 microns are inhaled into a patient&#39;s lungs and thereafter reach the patient&#39;s circulatory system.

CROSS-REFERENCES

This application is a continuation of Ser. No. 10/886,923 filed Jul. 7,2004 now U.S. Pat. No. 7,059,321 which is a continuation of applicationSer. No. 10/216,517 filed Aug. 9, 2002 now U.S. Pat. No. 6,792,940,which is a continuation of application Ser. No. 09/191,317, filed Nov.13, 1998 now abandoned, and which is a continuation-in-part ofapplication Ser. No. 09/171,518 filed Apr. 21, 1999 now U.S. Pat. No.6,119,953, which is a national phase filing of PCT/ES97/00034 under 35USC §371 filed Feb. 18, 1997 now WO 97/43048, which claims priority ofinternational application 9601101 filed May 13, 1996 under 35 U.S.C.§119 and also claims priority to P9702654 filed Dec. 17, 1997 under 35U.S.C. §119, all of which are incorporated herein by reference in theirentireties and to which applications we claim priority under 35 USC §120and 35 USC § 119.

FIELD OF THE INVENTION

This invention relates generally to the field of aerosols and moreparticularly to devices and methods for creating aerosols ofpharmaceutical formulations for delivery to a human patient, preferablyby inhalation.

BACKGROUND OF THE INVENTION

Aerosolizing formulations for inhalation has been considered as aconvenient alternative to injection for decades. This alternative toinjections is particularly interesting for drugs which cannot bedelivered orally, e.g. insulin. Although most compounds will effectivelymove from the lungs into the circulatory system there is considerableunpredictability in how much aerosolized formulation reaches the areasof the lungs where the material can move into the circulatory system.This results in inefficiency and unpredictability of dosing. A number ofdevices have been proposed for improving the efficiency of aerosoldelivery, monitoring patients and teaching patients to correctly usedelivery devices.

There are several different types of devices which use generallydifferent mechanisms and methodologies to produce aerosols forinhalation. The most commonly used device is a metered dose inhaler(MDI) which comprises a drug formulation container with the formulationincluding a low boiling point propellant. The formulation is held in thecontainer under pressure and a metered dose of formulation is releasedas an aerosol when the valve on the container is opened. The low boilingpoint propellant quickly evaporates or “flashes” when the formulation isexposed to atmospheric pressure outside the container. The particles offormulation containing the drug without the propellant are inhaled intothe patient's lungs and thereafter migrate into the patient'scirculatory system. There are a number of different types of MDIdevices. Devices of this type are disclosed in U.S. Pat. Nos. 5,404,871issued Apr. 11, 1995 and 5,364,838 issued Nov. 15, 1994.

Another type of device is the dry powder inhaler (DPI) device. Asindicated by the name such devices use formulations of dry powder whichpowder is blown into an aerosolized cloud via a burst of gas. TypicalDPI devices are shown in U.S. Pat. No. 5,775,320 issued Jul. 7, 1998 andU.S. Pat. No. 5,740,794 issued Apr. 21, 1998.

Yet another type of aerosol delivery device forces a formulation througha porous membrane. Formulation moving through the pores breaks up toform small particles which are inhaled by the patient. Devices of thistype are shown in U.S. Pat. No. 5,554,646 issued Aug. 13, 1996 and U.S.Pat. No. 5,522,385 issued Jun. 4, 1996.

Each of these devices has some advantages and disadvantages. The objectof each is substantially the same—to repeatedly produce a fine mistaerosol wherein the particles are substantially uniform in size andwithin a size range of about 1 micron to about 5 microns. A patient canbe accurately dosed if the device can repeatedly start with a givenamount of formula and produce a known amount of aerosol with particleshaving sizes within a known range. The present invention endeavors toprovide a device and method for obtaining accurate repeatable dosing ofa patient with an aerosol.

SUMMARY OF THE INVENTION

Aerosolized particles within a desired size range (e.g., 1 micron toabout 5 microns) are produced from a liquid formulation comprised of apharmaceutical active drug and a carrier. The particles produced allhave substantially the same particle diameter ±3% to ±30%, e.g. allparticles in the aerosol have a diameter of 2 microns ±3% to ±10%. Theformulation is provided in any desired manner (e.g., forced through achannel of a feeding needle and expelled out of an exit opening of theneedle). Simultaneously, gas contained in a pressure chamber (whichsurrounds at least the area where the formulation is provided, e.g.,surrounds the exit opening of the needle) is forced out of an openingpositioned in front of the formulation, e.g., directly in front of theflow path of the formulation being expelled from the feeding needle.Various parameters are adjusted to obtain a super critical flow ofliquid characterized by a stable liquid-gas interface and a stablecapillary jet of the liquid which forms particles on exiting the openingof the pressurized chamber which particles will all (90% or more) havesubstantially the same diameter, i.e., a monodisperse aerosol.

An object of the invention is to provide a device for aerosolizeddelivery of a pharmaceutically active drug formulation or a diagnosticformulation.

Another object is to provide a method of creating an aerosol ofconsistent particle size (±3 to 30% or preferably ±3 to 10% differencein diameter) which aerosol is inhalable by a patient for aerosolizeddelivery of drugs or diagnostics.

A feature of the invention is that the diameter of the opening fromwhich liquid is expelled, the diameter of the opening from which gas isexpelled and the distance between these two openings is adjustable andis adjusted to obtain a stable liquid-gas interface which results in astable capillary microjet being formed by the liquid expelled whichmicrojet is focused on an exit opening by the flow of surrounding gas.

Another feature of the invention is that the viscosities and velocitiesof the fluids can be chosen with consideration to other adjustedparameters to obtain a supercritical flow of liquid.

Another feature of the invention is that the liquid can be a singleliquid, two or more (miscible or immiscible) liquids mixed, a solutionor a suspension.

An advantage of the invention is that the gas flowing with the particlesprevents the particles from agglomerating thereby maintaining amonodisperse aerosol.

An advantage of the invention is that it consistently produces aerosolshaving particles with a desired particle diameter e.g. 1 to 5 microns.

An advantage of the invention is that the device of the invention isenergy efficient in terms of the energy used to create small particlesfor inhalation.

Another advantage is that the structure of the device and its use aresimple.

Another advantage is that clogging of the exit opening of the pressurechamber is substantially eliminated because liquid is kept out ofcontact with the surface of the exit opening by a surrounding focusedfunnel of gas which flows out of the pressure chamber exit opening.

Yet another advantage is that particles produced are substantiallysmaller in size than would be expected based on the diameter of the exitopening of the pressure chamber due to focusing the flow of the liquidwith the flow of surrounding gas.

An aspect of the invention is a hand-held, self-contained portable drugdelivery device which consistently produces small aerosolized particleswhich are relatively uniform in size.

Another aspect of the invention is a device and method which producesmultiple streams of aerosol thereby quickly aerosolizing a large dose offormulation for inhalation by a patient.

These and other aspects, objects, features and advantages will becomeapparent to those skilled in the art upon reading this disclosure incombination with the figures provided.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic view showing the basic components of oneembodiment of the invention with a cylindrical feeding needle as asource of formulation.

FIG. 2 is a schematic view of another embodiment of the invention withtwo concentric tubes as a source of formulation.

FIG. 3 is a schematic view of yet another embodiment showing awedge-shaped planar source of formulation. FIG. 3 a illustrates across-sectional side view of the planar feeding source and theinteraction of the fluids. FIG. 3 b show a frontal view of the openingsin the pressure chamber, with the multiple openings through which theatomizate exits the device. FIG. 3 c illustrates the channels that areoptionally formed within the planar feeding member. The channels arealigned with the openings in the pressure chamber.

FIG. 4 is a schematic view of a stable capillary microjet being formedand flowing through an exit opening to thereafter form a monodisperseaerosol.

FIG. 5 is a graph of data where 350 measured valves of d_(j)/d_(o)versus Q/Q_(o) are plotted.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Before the present aerosol device and method are described, it is to beunderstood that this invention is not limited to the particularcomponents and steps described, as such may, of course, vary. It is alsoto be understood that the terminology used herein is for the purpose ofdescribing particular embodiments only, and is not intended to belimiting, since the scope of the It must be noted that as used hereinand in the appended claims; the singular forms “a”, “and,” and “the”include plural referents unless the context clearly dictates otherwise.Thus, for example, reference to “a particle” includes a plurality ofparticles and reference to “a fluid ” includes reference to a mixture offluids, and equivalents thereof known to those skilled in the art, andso forth. present invention will be limited only by the appended claims.

Unless defined otherwise, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art to which this invention belongs. Although any methods andmaterials similar or equivalent to those described herein can be used inthe practice or testing of the present invention, the preferred methodsand materials are now described. All publications mentioned herein areincorporated herein by reference to disclose and describe the methodsand/or materials in connection with which the publications are cited.

The publications discussed herein are provided solely for theirdisclosure prior to the filing date of the present application. Nothingherein is to be construed as an admission that the present invention isnot entitled to antedate such publication by virtue of prior invention.Further, the dates of publication provided may be different from theactual publication dates which may need to be independently confirmed.

DEFINITIONS

The term “formulation” is used to describe any single liquid, mixture,solution, suspension or the like which is a flowable liquid at roomtemperature and comprises either a pharmaceutically active compound, adiagnostic compound or a compound which becomes an active compound afterentering a patient, e.g. prodrugs or DNA encoding proteins. Theformulation is preferably further comprised of a carrier and morepreferably is a liquid that has physical properties (e.g., viscosity)such that when the formulation is moved through a device of theinvention the formulation is aerosolized into particles (0.1 to 10microns in diameter) which are inhaled into the lungs of a patient andpreferably reach the circulatory system. The carrier may be anypharmaceutically acceptable material and is preferably a flowable liquidwhich is compatible with the active agent. Formulations are preferablysolutions, e.g., aqueous solutions, ethanolic solutions,aqueous/ethanolic solutions, saline solutions, colloidal suspensions andmicrocrystalline suspensions. Formulations can be solutions orsuspensions of drugs (including systemically active drugs, e.g. insulin,insulin analogs including monomeric insulin such as insulin lispro) inthe above mentioned liquids.

The term “carrier” shall mean a substantially inactive (biologically)component of a formulation such as a pharmaceutically acceptableexcipient material which an active ingredient such as a drug, adiagnostic agent, prodrug or a gene vector is mixed with, suspended ordissolved in. The carrier is preferably a flowable liquid. Usefulcarriers do not adversely interact with the drug, diagnostic or genevector and have properties which allow for the formation of aerosolizedparticles preferably particles having a diameter in the range of 0.1 to10.0 microns (more preferably 1 to 5 microns) when a formulationcomprising the carrier and active ingredient is aerosolized out of adevice of the invention. Carriers include water, ethanol, salinesolutions and mixtures thereof with pure water being preferred. Othercarriers can be used provided that they can be formulated to create asuitable aerosol and do not adversely effect the active component orhuman lung tissue.

The term “active compound” means a compound which is either apharmaceutically active drug, becomes or produces a therapeutic compoundin the body or is a detectably labeled compound. In turn, a“pharmaceutically active drug” or therapeutic compound shall beinterpreted to mean any pharmaceutically effective compound used in thetreatment of disease. A “detectably labeled compound” includes compoundswhich are radioactively labeled see U.S. Pat. No. 5,829,436 issued Nov.3, 1998 Re: detectable labels.

Pharmaceutical formulations for use in the invention may be comprised ofany therapeutic agent that can be atomized for patient delivery, e.g.for pulmonary delivery using an inhalation delivery system. Examples ofsuch pharmaceutical compositions may be found in the Physicians DeskReference (1998), and Remington: The Science and Practice of Pharmacy19^(th) Edition (1995), both of which are incorporated herein byreference. These formulations may be in any form capable of atomization,such as a solution, a suspension, an emulsion, a slurry, etc., providedthe dynamics of the form render it capable of forming a capillarymicrojet upon exposure to a second fluid.

The terms “electronic particle sensing device” and “electronic detectionsystem” refer to a device or other means for measuring and countingparticles present in an aerosol mist.

The term “reporting means” is used herein to indicate a means by whichinformation obtained from a monitoring device such as an electronicparticle sensing device or a device which monitors and recordsinformation relating to the patient's respiratory movement, may beprovided to the user.

The term “aerosol bolus” is used herein to describe a volume of airgreater than 50 ml and less than 4 liters which has suspended within itparticles of a formulation wherein the particles have a diameter in therange of 0.1 to 10 microns, preferably 1 to 5 microns, and preferablythe total volume of formulation is from 5 μl to 10,000 μl. About 10 μlof particles having a diameter of about 1 to 3 microns are present in avolume of about 50 ml to 2 liters, preferably 100 ml to 1,000 ml.

The term “aerosol” means any part of an aerosol bolus as describedabove.

The terms “air”, “particle free air”, “aerosol free air,” and the like,are used interchangeably herein to describe a volume of air which issubstantially free of other material and, in particular, free ofparticles intentionally added such as particles of formulation whichcreate the aerosol. The term means that the air does not includeparticles of formulation which have been intentionally added but is notintended to imply that the normal surrounding air has been filtered ortreated to remove all particles although filtering can take place. Airis the preferred gas to use with drug delivery it being noted that othernon-toxic gases, e.g., C0₂ can be used.

The term “measuring” describes an event whereby the (1) total lungcapacity, (2) inspiratory flow rate or (3) inspiratory volume of thepatient is measured and/or calculated. The information is preferablyused in order to determine an optimal point in the inspiratory cycle atwhich to release an aerosol and/or a particle free volume of air. Anactual measurement of both rate and volume may be made or the rate canbe directly measured and the volume calculated based on the measuredrate. The total lung capacity can be measured or calculated based on thepatient's height, sex and age. It is also preferable to continuemeasuring inspiratory flow during and after aerosol delivery and torecord inspiratory flow rate and volume before, during and after therelease of formulation. Such reading makes it possible to determine ifaerosolized formulation was properly delivered to the patient.

The term “monitoring” event shall comprise determining the character ofan aerosol such as the size, number and speed of the particles in theaerosol, further, monitoring may comprise measuring lung functions suchas inspiratory flow, inspiratory flow rate, and/or inspiratory volume sothat a patient's lung function as defined herein, can be evaluatedduring before and/or after delivery thereby making it possible toevaluate the effect of formulation such as respiratory drug delivery onthe patient's lung function.

The term “inspiratory flow rate” shall mean a value of air flow ratedetermined, calculated or measured based on the speed of the air passinga given point in a measuring device assuming atmospheric pressure ±5%and a temperature in the range of about 10° C. to 40° C.

The term “inspiratory flow” shall be interpreted to mean a value of airflow calculated based on the speed of the air passing a given pointalong with the volume of the air that has passed that point with thevolume calculation being based on integration of the flow rate data andassuming atmospheric pressure, ±5% and temperature in the range of about10° C. to about 40° C.

The term “inspiratory volume” shall mean a determined, measured orcalculated volume of air passing a given point into the lungs of apatient assuming atmospheric pressure ±5% and a temperature in the rangeof 10° C. to 40° C.

The terms “lung function” and “pulmonary function” are usedinterchangeably and shall be interpreted to mean physically measurableoperations of a lung including but not limited to (1) inspiratory and(2) expiratory flow rates as well as (3) lung volume, i.e., total lungcapacity. Methods of quantitatively determining pulmonary function areused to measure lung function. Quantitative determination of pulmonaryfunction may be important when delivering any formulation and inparticular respiratory drugs in order to direct the aerosol to aspecific area of the lung and to determine effectiveness. Methods ofmeasuring pulmonary function most commonly employed in clinical practiceinvolve timed measurement of inspiratory and expiratory maneuvers tomeasure specific parameters. For example, forced vital capacity (FVC)measures the total volume in liters exhaled by a patient forcefully froma deep initial inspiration. This parameter, when evaluated inconjunction with the forced expired volume in one second (FEV₁), allowsbronchoconstriction to be quantitatively evaluated. A problem withforced vital capacity determination is that the forced vital capacitymaneuver (i.e. forced exhalation from maximum inspiration to maximumexpiration) is largely technique dependent. In other words, a givenpatient may produce different FVC values during a sequence ofconsecutive FVC maneuvers. The FEF 25-75 or forced expiratory flowdetermined over the mid-portion of a forced exhalation maneuver tends tobe less technique dependent than the FVC. Similarly, the FEV₁ tends tobe less technique dependent than FVC. In addition to measuring volumesof exhaled air as indices of pulmonary function, the flow in liters perminute measured over differing portions of the expiratory cycle can beuseful in determining the status of a patient's pulmonary function. Inparticular, the peak expiratory flow, taken as the highest air flow ratein liters per minute during a forced maximal exhalation, is wellcorrelated with overall pulmonary function in a patient with asthma andother respiratory diseases. The present invention carries out treatmentby administering formulation in a delivery event and monitoring lungfunction in a monitoring event. A series of such events may be carriedout and repeated over time to determine if lung (see U.S. Pat. No.5,404,871) function is improved.

Each of the parameters discussed above is measured during quantitativespirometry. A patient's individual performance can be compared againsthis personal best data, individual indices can be compared with eachother for an individual patient (e.g. FEV₁ divided by FVC, producing adimensionless index useful in assessing the severity of acute asthmasymptoms), or each of these indices can be compared against an expectedvalue. Expected values for indices derived from quantitative spirometryare calculated as a function of the patient's sex, height, weight andage. For instance, standards exist for the calculation of expectedindices and these are frequently reported along with the actualparameters derived for an individual patient during a monitoring eventsuch as a quantitative spirometry test.

The terms “particles”, “aerosolized particles” and “aerosolizedparticles of formulation” are used interchangeably herein and shall meanparticles of formulation comprised of any pharmaceutically activeingredient, a compound which becomes such in the body (e.g. a prodrug orDNA encoding a protein) or a diagnostic compound. Any of these ispreferably with a carrier, (e.g., a pharmaceutically active respiratorydrug and carrier). Particles of a liquid formulation are formed uponforcing the formulation from anywhere it is provided (e.g., the feedingneedle) and then out of the pressure chamber exit orifice. The particleshave a size which is sufficiently small such that when the particles areformed they remain suspended in the air for a sufficient amount of timesuch that the patient can inhale the particles into the patient's lungs.The particles have a size in the range of 0.1 micron to about 10microns, preferably 1 to 5 microns. Particle diameter is an aerodynamicdiameter.

In referring to particle diameter, the diameter measurement given is the“aerodynamic diameter” measurement which is the diameter of a sphere ofunit density that has the same terminal sedimentation velocity in airunder normal atmospheric conditions as the particle in question. This ispointed out in that it is difficult to accurately measure the diameterof small particles using current technology and the shape may becontinually changing. Thus, the diameter of one particle of material ofa given density will be said to have the same diameter as anotherparticle of the same or a different material if the two particles havethe same terminal sedimentation velocity in air under the sameconditions.

DEVICE IN GENERAL

The present disclosure focuses on the use of an aerosol creation deviceto produce aerosolized liquid formulations for delivery of drugs topatients—preferably by inhalation into the lungs.

The basic technology of the invention comprises (1) a means forsupplying a first fluid; and (2) a pressure chamber supplied with asecond fluid. The first fluid is generally a liquid and preferablyaqueous. The second fluid is generally a gas and preferably air.However, the first fluid may be a gas and second fluid a liquid or bothfluids may be liquid provided the first and second fluid aresufficiently different from each other (immiscible) so as to allow forthe formation of a stable microjet of the first fluid moving from thesupply means to an exit port of the pressure chamber. Notwithstandingthese different combinations of gas-liquid, liquid-gas, andliquid-liquid the invention is generally described with a liquidformulation being expelled from the supply means and forming a stablemicrojet due to interaction with surrounding air flow focusing themicrojet to flow out of an exit of the pressure chamber.

Formation of the microjet and its acceleration and ultimate particleformation are based on the abrupt pressure drop associated with thesteep acceleration experienced by the first fluid (e.g., a liquid) onpassing through an exit orifice of the pressure chamber which holds thesecond fluid. On leaving the chamber the flow undergoes a large pressuredifference between the first fluid (e.g., a liquid) and the second fluid(e.g., a gas), which in turn produces a highly curved zone on the firstfluid (e.g., liquid) surface near the exit port of the pressure chamberand in the formation of a cuspidal point from which a steady microjetflows provided the amount of the first fluid (e.g., the liquid)withdrawn through the exit port of the pressure chamber is replenished.Thus, in the same way that a glass lens or a lens of the eye focuseslight to a given point, the flow of the gas surrounds and focuses theliquid into a stable microjet. The focusing effect of the surroundingflow of gas creates a stream of liquid which is substantially smaller indiameter than the diameter of the exit orifice of the pressure chamber.This allows liquid to flow out of the pressure chamber orifice withouttouching the orifice, providing advantages including (1) clogging of theexit orifice is virtually eliminated, (2) contamination of flow due tocontact with substances (e.g. bacteria) on the orifice opening isvirtually eliminated, and (3) the diameter of the stream and theresulting particles are smaller than the diameter of the exit orifice ofthe chamber. This is particularly desirable because it is difficult toprecisely engineer holes which are very small in diameter. Further, inthe absence of the focusing effect (and formation a stable microjet)flow of liquid out of an opening will result in particles which haveabout twice the diameter of the exit opening. An additional advantage isthat the particles are not prone to agglomeration following exit fromthe chamber.

Specific embodiments of aerosol creation devices are now described.

EMBODIMENT OF FIG. 1

A first embodiment of the invention where the supply means is acylindrical feeding needle supplying liquid into a pressurized chamberof gas is described below with reference to FIG. 1.

The components of the embodiment of FIG. 1 are as follows:

1. Feeding needle—also referred to generally as a fluid source and atube.

2. End of the feeding needle used to insert the liquid to be atomized.

3. Pressure chamber.

4. Orifice used as gas inlet.

5. End of the feeding needle used to evacuate the liquid to be atomized.

6. Orifice through which withdrawal takes place.

7. Atomizate (spray)—also referred to as aerosol.

D₀=diameter of the feeding needle; d₀=diameter of the orifice throughwhich the microjet is passed; e=axial length of the orifice throughwhich withdrawal takes place; H=distance from the feeding needle to themicrojet outlet; P₀=pressure inside the chamber; P_(α)=atmosphericpressure.

An aerosolized drug delivery device of the invention may be any size butis preferably designed to be a hand-held, portable, self-containeddevice weighing less than 1 kilogram. Although the device can beconfigured in a variety of designs, the different designs will allinclude the essential components shown in FIG. 1 or components whichperform an equivalent function and obtain the desired results.Specifically, a drug delivery device of the invention will be comprisedof at least one source of formulation (e.g., a feeding needle with anopening 2) into which a liquid flowable formulation can be fed and anexit opening 5 from which the formulation can be expelled. The feedingneedle 1, or at least its exit opening 5, is encompassed by a pressurechamber 3. The chamber 3 has inlet opening 4 which is used to feed gasinto the chamber 3 and an exit opening 6 through which gas from thepressure chamber and liquid formulation from the feeding needle 3 areexpelled creating an aerosol.

In FIG. 1, the feeding needle and pressure chamber are configured toobtain a desired result of producing an aerosol wherein the particlesare small and uniform in size. Preferably the particles have a sizewhich is in a range of 0.1 to 10 microns, more preferably 1 to 5microns. Particles of less than 1 micron in diameter can be produced viathe present invention. However, particles below 1 micron may be toosmall for inhalation as the particles may not settle in the lung duringa normal breath hold and as such would be exhaled. The particles of anygiven aerosol all have about the same diameter with a relative standarddeviation of 10% to 30% or more preferably 3% to 20%. Stating thatparticles of the aerosol have a particle diameter in a range of 1 to 5microns does not mean that different particles will have differentdiameters and that some will have a diameter of 1 micron while others of5 microns. The particles in a given aerosol will all (preferably about90% or more) have the same diameter ±3% to ±30%. For example, theparticles of a given aerosol will have a diameter of 2 microns ±3% to±10%.

Such a monodisperse aerosol is created using the components andconfiguration as described above. However, other components andconfigurations will occur to those skilled in the art. The object ofeach design will be to supply formulation so that it creates a stablecapillary microjet which is accelerated and stabilized by tangentialviscous stress exerted by the gas on the liquid surface. The stablemicrojet created by the gas leaves the area of the pressurized gas(e.g., leaves the pressure chamber and exits the pressure chamberorifice) and splits into particles which have the desired size anduniformity.

The aerosol created is a monodisperse aerosol meaning that the size ofthe particles produced are relatively uniform in size. The relativestandard deviation in particle size is in the range of from about 10% toabout 30%, preferably 3% to 10% and most preferably 3% or less. The sizeof aerosolized particles useful for inhalation is a diameter in therange of from about 0.1 micron to about 10 microns, more preferablyabout 1 micron to about 3 microns.

For purposes of simplicity the remainder of the detailed description ofthe operation of the device of FIG. 1 will refer to the first fluid asliquid and the second fluid as gas. The parameter window used (i.e. theset of special values for the liquid properties, flow-rate used, feedingneedle diameter, orifice diameter, pressure ratio, etc.) should be largeenough to be compatible with virtually any liquid (dynamic viscositiesin the range from 10⁻⁴ to 1 kg m⁻¹s⁻¹); in this way, the capillarymicrojet that emerges from the end of the feeding needle is absolutelystable and perturbations produced by breakage of the jet cannot travelupstream. Downstream, the microjet splits into evenly shaped dropssimply by effect of capillary instability (see, for example, Rayleigh,“On the instability of jets”, Proc. London Math. Soc., 4-13, 1878),similar in a manner to a laminar capillary jet falling from a half-opentap.

When the stationary, steady interface is created, the capillary jet thatemerges from the end of the drop at the outlet of the feeding point isconcentrically withdrawn into the nozzle. After the jet emerges from thedrop, the liquid is accelerated by tangential sweeping forces exerted bythe gas stream flowing on its surface, which gradually decreases the jetcross-section. Stated differently the gas flow acts as a lens andfocuses and stabilizes the microjet as it moves toward and into the exitorifice of the pressure chamber.

The forces exerted by the second fluid (e.g., a gas) flow on the firstfluid (e.g., a liquid) surface should be steady enough to preventirregular surface oscillations. Therefore, any turbulence in the gasmotion should be avoided; even if the gas velocity is high, thecharacteristic size of the orifice should ensure that the gas motion islaminar (similar to the boundary layers formed on the jet and on theinner surface of the nozzle or hole).

STABLE CAPILLARY MICROJET

FIG. 4 illustrates the interaction of a liquid and a gas to formatomizate using the method of the invention. The feeding needle 60 has acircular exit opening 61 with an internal radius R₀ which feeds a liquid62 out of the end, forming a drop with a radius in the range of R₀ to R₀plus the thickness of the wall of the needle. The exiting liquid formsan infinite amount of liquid streamlines 63 that interact with thesurrounding gas to form a stable cusp at the interface 64 of the twofluids. The surrounding gas also forms an infinite number of gasstreamlines 65, which interact with the exiting liquid to create avirtual focusing funnel 66. The exiting liquid is focused by thefocusing funnel 66 resulting in a stable capillary microjet 67, whichremains stable until it exits the opening 68 of the pressure chamber 69.After exiting the pressure chamber, the microjet begins to break-up,forming monodispersed particles 70.

The gas flow, which affects the liquid withdrawal and its subsequentacceleration after the jet is formed, should be very rapid but alsouniform in order to avoid perturbing the fragile capillary interface(the surface of the drop that emerges from the jet).

Liquid flows out of the end of a capillary tube and forms a small liquiddrop at the end. The tube has an internal radius R_(o). The drop has aradius in a range of from R_(o) to R_(o) plus the structural thicknessof the tube as the drop exits the tube, and thereafter the drop narrowsin circumference to a much smaller circumference as is shown in theexpanded view of the tube (i.e. feeding needle) 5 as shown in FIGS. 1and 4.

As illustrated in FIG. 4, the exit opening 61 of the capillary tube 60is positioned close to an exit opening 68 in a planar surface of apressure chamber 69. The exit opening 68 has a minimum diameter D and isin a planar member with a thickness L. The diameter D is referred to asa minimum diameter because the opening may have a conical configurationwith the narrower end of the cone positioned closer to the source ofliquid flow. Thus, the exit opening may be a funnel-shaped nozzlealthough other opening configurations are also possible, e.g. an hourglass configuration. Gas in the pressure chamber continuously flows outof the exit opening. The flow of the gas causes the liquid drop expelledfrom the tube to decrease in circumference as the liquid moves away fromthe end of the tube in a direction toward the exit opening of thepressure chamber.

In actual use, it can be understood that the opening shape whichprovokes maximum gas acceleration (and consequently the most stable cuspand microjet with a given set of parameters) is a conically shapedopening in the pressure chamber. The conical opening is positioned withits narrower end toward the source of liquid flow.

The distance between the end 61 of the tube 60 and the beginning of theexit opening 68 is H. At this point it is noted that R_(o), D, H and Lare all preferably on the order of hundreds of microns. For example,R_(o)=400 μm, D=150 μm, H=1 mm, L=300 μm. However, each could be 1/100to 100× these sizes.

The end of the liquid stream develops a cusp-like shape at a criticaldistance from the exit opening 68 in the pressure chamber 69 when theapplied pressure drop ΔP_(g) across the exit opening 68 overcomes theliquid-gas surface tension stresses γ/R* appearing at the point ofmaximum curvature—e.g. 1/R* from the exit opening.

A steady state is then established if the liquid flow rate Q ejectedfrom the drop cusp is steadily supplied from the capillary tube. This isthe stable capillary cusp which is an essential characteristic of theinvention needed to form the stable microjet. More particularly, asteady, thin liquid jet with a typical diameter d_(j) is smoothlyemitted from the stable cusp-like drop shape and this thin liquid jetextends over a distance in the range of microns to millimeters. Thelength of the stable microjet will vary from very short (e.g. 1 micron)to very long (e.g. 50 mm) with the length depending on the (1) flow-rateof the liquid and (2) the Reynolds number of the gas stream flowing outof the exit opening of the pressure chamber. The liquid jet is thestable capillary microjet obtained when supercritical flow is reached.This jet demonstrates a robust behavior provided that the pressure dropΔP_(g) applied to the gas is sufficiently large compared to the maximumsurface tension stress (on the order of γ/d_(j)) that act at theliquid-gas interface. The jet has a slightly parabolic axial velocityprofile which is, in large part, responsible for the stability of themicrojet. The stable microjet is formed without the need for otherforces, i.e. without adding force such as electrical forces on a chargedfluid. However, for some applications it is preferable to add charge toparticles, e.g. to cause the particles to adhere to a given surface. Theshaping of liquid exiting the capillary tube by the gas flow forming afocusing funnel creates a cusp-like meniscus resulting in the stablemicrojet. This is a fundamental characteristic of the invention.

The fluid stream flowing from the tube has substantially more densityand develops substantially more inertia as compared to the gas, whichhas lower viscosity than the liquid. These characteristics contribute tothe formation of the stable capillary jet. The stable capillary microjetis maintained stably for a significant distance in the direction of flowaway from the exit from the tube. The liquid is, at this point,undergoing “supercritical flow.” The microjet eventually destabilizesdue to the effect of surface tension forces. Destabilization resultsfrom small natural perturbations moving downstream, with the fastestgrowing perturbations being those which govern the break up of themicrojet, eventually creating a uniform sized monodisperse aerosol 70 asshown in FIG. 4.

The microjet, even as it initially destabilizes, passes out of the exitorifice of the pressure chamber without touching the peripheral surfaceof the exit opening. This provides an important advantage of theinvention which is that the exit opening 68 (which could be referred toas a nozzle) will not clog from residue and/or deposits of the liquid.Clogging is a major problem with very small nozzles and is generallydealt with by cleaning or replacing the nozzle. When fluid contacts thesurfaces of a nozzle opening some fluid will remain in contact with thenozzle when the flow of fluid is shut off. The liquid remaining on thenozzle surface evaporates leaving a residue. After many uses over timethe residue builds up and clogging takes place. The present inventionsubstantially reduces or eliminates this clogging problem.

MATHEMATICS OF A STABLE MICROJET

Cylindrical coordinates (r,z) are chosen for making a mathematicalanalysis of a stable microjet, i.e. liquid undergoing “supercriticalflow.” The cusp-like meniscus formed by the liquid coming out of thetube is pulled toward the exit of the pressure chamber by a pressuregradient created by the flow of gas.

The cusp-like meniscus formed at the tube's mouth is pulled towards thehole by the pressure gradient created by the gas stream. From the cuspof this meniscus, a steady liquid thread with the shape of radius r=ξ iswithdrawn through the hole by the action of both the suction effect dueto ΔP_(g), and the tangential viscous stresses τ_(s) exerted by the gason the jet's surface in the axial direction. The averaged momentumequation for this

$\begin{matrix}{{{\frac{\mathbb{d}\;}{\mathbb{d}_{z}}\left\lbrack {P_{1} + \frac{\rho_{1}Q^{2}}{2{\prod^{2}\xi^{4}}}} \right\rbrack} = \frac{2_{\tau_{s}}}{\xi}},} & (1)\end{matrix}$configuration may be written:

where Q is the liquid flow rate upon exiting the feeding tube, P₁ is theliquid pressure, and ρ₁ is the liquid density, assuming that the viscousextensional term is negligible compared to the kinetic energy term, aswill be subsequently justified. In addition, liquid evaporation effectsare neglected. The liquid pressure P₁ is given by the capillaryequation.P ₁ =P ₈+γ/ξ.  (2)

where γ is the liquid-gas surface tension. As shown in the Examples, thepressure drop ΔP_(g) is sufficiently large as compared to the surfacetension stress γ/ξ to justify neglecting the latter in the analysis.This scenario holds for the whole range of flow rates in which themicrojet is absolutely stable. In fact, it will be shown that, for agiven pressure drop ΔP_(g), the minimum liquid flow rate that can besprayed in steady jet conditions is achieved when the surface tensionstress γ/ξ is of the order of the kinetic energy of the liquidρ₁Q²/(2π²ξ⁴), since the surface tension acts like a “resistance” to themotion (it appears as a negative term in the right-hand side term of Eq.(1)). Thus,

Q min - ( γ ⁢ ⁢ d j 3 ρ 1 ) ( 3 )

For sufficiently large flow rates Q compared to Q_(min), the simplifiedaveraged momentum equation in the axial direction can be expressed as

$\begin{matrix}{{{\frac{\mathbb{d}\;}{\mathbb{d}_{z}}\left( \frac{\rho_{1}Q^{2}}{2{\prod^{2}\xi^{4}}} \right)} = {\frac{\mathbb{d}P_{g}}{\mathbb{d}_{z}} + \frac{2_{\tau_{s}}}{\xi}}},} & (4)\end{matrix}$

where one can identify the two driving forces for the liquid flow on theright-hand side. This equation can be integrated provided the followingsimplification is made: if one uses a thin plate with thickness L of theorder or smaller than the hole's diameter D (which minimizes downstreamperturbations in the gas flow), the pressure gradient up to the holeexit is on the average much larger than the viscous shear term 2τ_(s)/ξowning to the surface stress. On the other hand, the axial viscous termis of the order O[μ²Q/D²d_(j) ²], since the hole diameter D is actuallythe characteristic distance associated with the gas flow at the hole'sentrance in both the radial and axial directions. This term is verysmall compared to the pressure gradient in real situations, providedthat ΔP_(g)>>μ²/D²ρ₁ (which holds, e.g., for liquids with viscosities aslarge as 100 cpoises, using hole diameters and pressure drops as smallas D˜10 μm and ΔP_(g)≧100 mbar). The neglect of all viscous terms in Eq.(4) is then justified. Notice that in this limit on the liquid flow isquasi-isentropic in the average (the liquid almost follows Bernoulliequation) as opposed to most micrometric extensional flows. Thus,integrating (4) from the stagnation regions of both fluids up to theexit, one obtains a simple and universal expression for the jet diameterat the hole exit:

d j - ( 8 ⁢ ⁢ ρ 1 ∏ 2 ⁢ Δ ⁢ ⁢ P g ) 1 / 4 ⁢ Q , ( 5 )

which for a given pressure drop ΔP_(g) is independent of geometricalparameters (hole and tube diameters, tube-hole distance, etc.), liquidand gas viscosities, and liquid-gas surface tension. This diameterremains almost constant up to the breakup point since the gas pressureafter the exit remains constant.

MONODISPERSE PARTICLES

Above the stable microjet undergoing “supercritical flow” is describedand it can be seen how this aspect of the invention can be made use ofin a variety of industrial applications—particularly where the flow ofliquid through small holes creates a clogging problem. An equallyimportant aspect of the invention is obtained after the microjet leavesthe pressure chamber.

When the microjet exits the pressure chamber the liquid pressure P₁becomes (like the gas pressure P_(g)) almost constant in the axialdirection, and the jet diameter remains almost constant up to the pointwhere it breaks up by capillary instability. Defining a Weber numberWe=(ρ_(g)v_(g) ²d_(j))/γ≅2ΔP_(g)d_(j)/γ (where v_(g) is the gas velocitymeasured at the orifice), below a certain experimental value We_(c)˜40the breakup mode is axisymmetric and the resulting droplet stream ischaracterized by its monodispersity provided that the fluctuations ofthe gas flow do not contribute to droplet coalescence (thesefluctuations occur when the gas stream reaches a fully developedturbulent profile around the liquid jet breakup region). Above thisWe_(c) value, sinuous nonaxisymmetric disturbances, coupled to theaxisymmetric ones, become apparent. For larger We numbers, the nonlineargrowth rate of the sinuous disturbances seems to overcome that of theaxisymmetric disturbances. The resulting spray shows significantpolydispersity in this case. Thus, it can be seen that by controllingparameters to keep the resulting Weber number to 40 or less, allows theparticles formed to be all substantially the same size. The sizevariation is about ±3% to ±30% and move preferably ±3% to ±10%. Theseparticles can have a desired size e.g. 0.1 microns to 50 microns.

The shed vorticity influences the breakup of the jet and thus theformation of the particles. Upstream from the hole exit, in theaccelerating region, the gas stream is laminar. Typical values of theReynolds number range from 500 to 6000 if a velocity of the order of thespeed of sound is taken as characteristic of the velocity of the gas.Downstream from the hole exit, the cylindrical mixing layer between thegas stream and the stagnant gas becomes unstable by the classicalKelvin-Helmholtz instability. The growth rate of the thickness of thislayer depends on the Reynolds number of the flow and ring vortices areformed at a frequency of the order of v_(g)/D, where D is the holediameter. Typical values of v_(g) and D as those found in ourexperimental technique lead to frequencies or the order of MHz which arecomparable to the frequency of drop production (of order of t_(b) ⁻¹).

Given the liquid flow rate and the hole diameter, a resonance frequencywhich depends on the gas velocity (or pressure difference driving thegas stream) can be adjusted (tuned) in such a way that vortices act as aforcing system to excite perturbations of a determined wavelength on thejet surface. Experimental results obtained clearly illustrates thedifferent degree of coupling between the two gas-liquid coaxial jets. Inone set of experimental results the particle sizes are shown to have aparticle size of about 5.7 microns with a standard deviation of 12%.This results when the velocity of the gas has been properly tuned tominimize the dispersion in the size of droplets resulting from the jetbreakup. In this case, the flow rate of the liquid jet and its diameterare 0.08 μl s⁻¹ and 3 μm, respectively. Data have been collected using aMASTERSIZER from MALVERN Instruments. As the degree of couplingdecreases, perturbations at the jet surface of different wavelengthsbecome excited and, as it can be observed from the size distributions,the dispersion of the spray increases.

It is highly desirable in a number of different industrial applicationsto have particles which are uniform in size or to create aerosols ofliquid particles which are uniform in size. For example, particles of aliquid formation containing a pharmaceutically active drug could becreated and designed to have a diameter of about 2 microns ±3%. Theseparticles could be inhaled into the lungs of a patient forintrapulmonary drug delivery. Moreover, particle size can be adjusted totarget a particular area of the respiratory tract.

The gas flow should be laminar in order to avoid a turbulentregime—turbulent fluctuations in the gas flow which have a highfrequency and would perturb the liquid-gas interface. The Reynoldsnumbers reached at the orifice are

${Re} = {\frac{v_{g}d_{0}}{v_{g}} - 4000}$

where v_(g) is the kinematic viscosity of the gas. Even though thisnumber is quite high, there are large pressure gradients downstream (ahighly convergent geometry), so that a turbulent regime is very unlikelyto develop.

The essential difference from existing pneumatic atomizers (whichpossess large Weber numbers) and the present invention is that the aimof the present invention is not to rupture the liquid-gas interface butthe opposite, i.e. to increase the stability of the interface until acapillary jet is obtained. The jet, which will be very thin provided thepressure drop resulting from withdrawal is high enough, splits intodrops the sizes of which are much more uniform than those resulting fromdisorderly breakage of the liquid-gas interface in existing pneumaticatomizers.

The proposed atomization system obviously requires delivery of theliquid to be atomized and the gas to be used in the resulting spray.Both should be fed at a rate ensuring that the system lies within thestable parameter window. Multiplexing is effective when the flow-ratesneeded exceed those on an individual cell. More specifically, aplurality of feeding sources or feeding needles may be used to increasethe rate at which aerosols are created. The flow-rates used should alsoensure the mass ratio between the flows is compatible with thespecifications of each application.

The gas and liquid can be dispensed by any type of continuous deliverysystem (e.g. a compressor or a pressurized tank the former and avolumetric pump or a pressurized bottle the latter). If multiplexing isneeded, the liquid flow-rate should be as uniform as possible amongcells; this may entail propulsion through several capillary needles,porous media or any other medium capable of distributing a uniform flowamong different feeding points.

Each individual atomization device should consist of a feeding point (acapillary needle, a point with an open microchannel, a microprotuberanceon a continuous edge, etc.) 0.002-2 mm (but, preferentially 0.01-0.4 mm)in diameter, where the drop emerging from the microjet can be anchored,and a small orifice 0.002-2 mm (preferentially 0.01-0.25 mm) in diameterfacing the drop and separated 0.01-2 mm (preferentially 0.2-0.5 mm) fromthe feeding point. The orifice communicates the withdrawal gas aroundthe drop, at an increased pressure, with the zone where the atomizate isproduced, at a decreased pressure. The atomizer can be made from avariety of materials (metal, polymers, ceramics, glass).

FIG. 1 depicts a tested prototype where the liquid to be atomized isinserted through one end of the system 2 and the propelling gas inintroduced via the special inlet 4 in the pressure chamber 3. Theprototype was tested at gas feeding rates from 100 to 2000 mBar abovethe atmospheric pressure P_(α)at which the atomized liquid wasdischarged. The whole enclosure around the feeding needle 1 was at apressure P₀>P_(α). The liquid feeding pressure, P₁, should always beslightly higher than the gas propelling pressure, P₀. Depending on thepressure drop in the needle and the liquid feeding system, the pressuredifference (P₁-P₀>0) and the flow-rate of the liquid to be atomized, Q,are linearly related provided the flow is laminar—which is indeed thecase with this prototype. The critical dimensions are the distance fromthe needle to the plate (H), the needle diameter (D₀), the diameter ofthe orifice through which the microjet 6 is discharged (d₀) and theaxial length, e, of the orifice (i.e. the thickness of the plate wherethe orifice is made). In this prototype, H was varied from 0.3 to 0.7 mmon constancy of the distances (D₀=0.45 mm, d₀-0.2 mm) and e-0.5 mm. Thequality of the resulting spray 7 did not vary appreciably with changesin H provided the operating regime (i.e. stationary drop and microjet)was maintained. However, the system stability suffered at the longer Hdistances (about 0.7 mm). The other atomizer dimensions had no effect onthe spray or the prototype functioning provided the zone around theneedle (its diameter) was large enough relative to the feeding needle.

WEBER NUMBER

Adjusting parameters to obtain a stable capillary microjet and controlits breakup into monodisperse particle is governed by the Weber numberand the liquid-to-gas velocity ratio or α which equal V₁/V_(g). TheWeber number or “We” is defined by the following equation:

${We} = \frac{\rho_{g}V_{g}^{2}d}{\gamma}$

wherein ρ_(g) is the density of the gas, d is the diameter of the stablemicrojet, γ is the liquid-gas surface tension, and V_(g) ² is thevelocity of the gas squared.

When carrying out the invention the parameters should be adjusted sothat the Weber number is greater than 1 in order to produce a stablecapillary microjet. However, to obtain a particle dispersion which ismonodisperse (i.e. each particle has the same size ±3 to ±30%) theparameters should be adjusted so that the Weber number is less than 40.The monodisperse aerosol is obtained with a Weber number in a range ofabout 1 to about 40 when the breaking time is sufficiently small toavoid non-symmetric perturbations (1≦We≦40).

OHNESORGE NUMBER

A measure of the relative importance of viscosity on the jet breakup canbe estimated from the Ohnesorge number defined as the ratio between twocharacteristic times: the viscous time t_(v) and the breaking timet_(b). The breaking time t_(b) is given by [see Rayleigh (1878)]

t b - ( ρ l ⁢ d 2 γ ) . ( 2 )

Perturbations on the jet surface are propagated inside by viscousdiffusion in times t_(v) of the order oft_(v—)ρ₁d²_μ₁,  (3)

where μ₁ is the viscosity of the liquid. Then, the Ohnesorge number, Oh,results

Oh = μ 1 ( ρ l ⁢ γ ⁢ ⁢ d ) . ( 4 )

If this ratio is much smaller than unity viscosity plays no essentialrole in the phenomenon under consideration. Since the maximum value ofthe Ohnesorge number in actual experiments conducted is as low as3.7×10⁻², viscosity plays no essential role during the process of jetbreakup.

EMBODIMENT OF FIG. 2

A variety of configurations of components and types of fluids willbecome apparent to those skilled in the art upon reading thisdisclosure. These configurations and fluids are encompassed by thepresent invention provided they can produce a stable capillary microjetof a first fluid from a source to an exit port of a pressure chambercontaining a second fluid. The stable microjet is formed by the firstfluid flowing from the feeding source to the exit port of the pressurechamber being accelerated and stabilized by tangential viscous stressexerted by the second fluid in the pressure chamber on the surface ofthe first fluid forming the microjet. The second fluid forms a focusingfunnel when a variety of parameters are correctly tuned or adjusted. Forexample, the speed, pressure, viscosity and miscibility of the first andsecond fluids are chosen to obtain the desired results of a stablemicrojet of the first fluid focused into the center of a funnel formedwith the second fluid. These results are also obtained by adjusting ortuning physical parameters of the device, including the size of theopening from which the first fluid flows, the size of the opening fromwhich both fluids exit, and the distance between these two openings.

The embodiment of FIG. 1 can, itself, be arranged in a variety ofconfigurations. Further, as indicated above, the embodiment may includea plurality of feeding needles.

A plurality of feeding needles may be configured concentrically in asingle construct, as shown in FIG. 2.

The components of the embodiment of FIG. 2 are as follows:

21. Feeding needle—tube or source of fluid.

22. End of the feeding needle used to insert the liquids to be atomized.

23. Pressure chamber.

24. Orifice used as gas inlet.

25. End of the feeding needle used to evacuate the liquid to beatomized.

26. Orifice through which withdrawal takes place.

27. Atomizate (spray) or aerosol.

28. First liquid to be atomized (inner core of particle).

29. Second liquid to be atomized (outer coating of particle).

30. Gas for creation of microjet.

31. Internal tube of feeding needle.

32. External tube of feeding needle.

D=diameter of the feeding needle; d=diameter of the orifice throughwhich the microjet is passed; e=axial length of the orifice throughwhich withdrawal takes place; H=distance from the feeding needle to themicrojet outlet; γ=surface tension; P₀=pressure inside the chamber;P_(α)=atmospheric pressure.

The embodiment of FIG. 2 is preferably used when attempting to form aspherical particle of one substance coated by another substance. Thedevice of FIG. 2 is comprised of the same basic component as per thedevice of FIG. 1 and further includes a second feeding source 32 whichis positioned concentrically around the first cylindrical feeding source31. The second feeding source may be surrounded by one or moreadditional feeding sources with each concentrically positioned aroundthe preceding source. The outer coating may be used for a variety ofpurposes, including: coating particles to prevent small particles fromsticking together; to obtain a sustained release effect of the activecompound (e.g. a pharmaceutically active drug) inside, and/or to maskflavors; and to protect the stability of another compound (e.g. apharmaceutically active drug) contained therein.

The process is based on the microsuction which the liquid-gas orliquid-liquid interphase undergoes (if both are immiscible), when saidinterphase approaches a point beginning from which one of the fluids issuctioned off while the combined suction of the two fluids is produced.The interaction causes the fluid physically surrounded by the other toform a capillary microjet which finally breaks into spherical drops. Ifinstead of two fluids (gas-liquid), three or more are used that flow ina concentric manner by injection using concentric tubes, a capillary jetcomposed of two or more layers of different fluids is formed which, whenit breaks, gives rise to the formation of spheres composed of severalapproximately concentric spherical layers of different fluids. The sizeof the outer sphere (its thickness) and the size of the inner sphere(its volume) can be precisely adjusted. This can allow the manufactureof coated particles for a variety of end uses. For example the thicknessof the coating can be varied in different manufacturing events to obtaincoated particles which have gradually decreasing thicknesses to obtain acontrolled release effect of the contents, e.g. a pharmaceuticallyactive drug. The coating could merely prevent the particles fromdegrading, reacting, or sticking together.

The method is based on the breaking of a capillary microjet composed ofa nucleus of one liquid or gas and surrounded by another or otherliquids and gases which are in a concentric manner injected by a specialinjection head, in such a way that they form a stable capillary microjetand that they do not mix by diffusion during the time between when themicrojet is formed and when it is broken. When the capillary microjet isbroken into spherical drops under the proper operating conditions, whichwill be described in detail below, these drops exhibit a sphericalnucleus, the size and eccentricity of which can be controlled.

In the case of spheres containing two materials, the injection head 25consists of two concentric tubes with an external diameter on the orderof one millimeter. Through the internal tube 31 is injected the materialthat will constitute the nucleus of the microsphere, while between theinternal tube 31 and the external tube 32 the coating is injected. Thefluid of the external tube 32 joins with the fluid of tube 31 as thefluids exit the feeding needle, and the fluids (normally liquids) thusinjected are accelerated by a stream of gas that passes through a smallorifice 24 facing the end of the injection tubes. When the drop inpressure across the orifice 24 is sufficient, the liquids form acompletely stationary capillary microjet, if the quantities of liquidsthat are injected are stationary. This microjet does not touch the wallsof the orifice, but passes through it wrapped in the stream of gas orfunnel formed by gas from the tube 32. Because the funnel of gas focusesthe liquid, the size of the exit orifice 26 does not dictate the size ofthe particles formed.

When the parameters are correctly adjusted, the movement of the liquidis uniform at the exit of the orifice 26 and the viscosity forces aresufficiently small so as not to alter either the flow or the propertiesof the liquids; for example, if there are biochemical molecularspecimens having a certain complexity and fragility, the viscous forcesthat would appear in association with the flow through a micro-orificemight degrade these substances.

FIG. 2 shows a simplified diagram of the feeding needle 21, which iscomprised of the concentric tubes 30, 31 through the internal andexternal flows of the fluids 28, 29 that are going to compose themicrospheres comprised of two immiscible fluids. The difference inpressures P₀-P_(α) (P₀>P_(α)) through the orifice 26 establishes a flowof gas present in the chamber 23 and which is going to surround themicrojet at its exit. The same pressure gradient that moves the gas isthe one that moves the microjet in an axial direction through the hole26, provided that the difference in pressures P₀-P_(α) is sufficientlygreat in comparison with the forces of surface tension, which create anadverse gradient in the direction of the movement.

There are two limitations for the minimum sizes of the inside andoutside jets that are dependent (a) on the surface tensions γ1 of theoutside liquid 29 with the gas 30 and γ2 of the outside liquid 29 withthe inside liquid 28, and (b) on the difference in pressures ΔP=P₀-P_(α)through the orifice 26. In the first place, the jump in pressures ΔPmust be sufficiently great so that the adverse effects of the surfacetension are minimized. This, however, is attained for very modestpressure increases: for example, for a 10 micron jet of a liquid havinga surface tension of 0.05 N/m (tap water), the necessary minimum jump inpressure is in the order of 0.05 (N/m)/0.00001 m=ΔP=50 mBar. But, inaddition, the breakage of the microjet must be regular andaxilsymmetric, so that the drops will have a uniform size, while theextra pressure ΔP cannot be greater than a certain value that isdependent on the surface tension of the outside liquid with the gas γland on the outside diameter of the microjet. It has been experimentallyshown that this difference in pressures cannot be greater than 20 timesthe surface tension γl divided by the outside radius of the microjet.

Therefore, given some inside and outside diameters of the microjet,there is a range of operating pressures between a minimum and a maximum;nonetheless, experimentally the best results are obtained for pressuresin the order of two to three times the minimum.

The viscosity values of the liquids must be such that the liquid withthe greater viscosity μ_(max) verifies, for a diameter d of the jetpredicted for this liquid and a difference through the orifice ΔP, theinequality:

$\mu_{MAX} \leq \frac{\Delta\;{Pd}^{2}D}{Q}$

With this, the pressure gradients can overcome the extensional forces ofviscous resistance exerted by the liquid when it is suctioned toward theorifice.

Moreover, the liquids must have very similar densities in order toachieve the concentricity of the nucleus of the microsphere, since therelation of velocities between the liquids moves according to the squareroot of the densities v1/v2=(ρ2/ρ1)1/2 and both jets, the inside jet andthe outside jet, must assume the most symmetrical configurationpossible, which does not occur if the liquids have different velocities(FIG. 2). Nonetheless, it has been experimentally demonstrated that, onaccount of the surface tension γ2 between the two liquids, the nucleustends to migrate toward the center of the microsphere, within prescribedparameters.

When two liquids and gas are used on the outside, the distance betweenthe planes of the mouths of the concentric tubes can vary, without thecharacteristics of the jet being substantially altered, provided thatthe internal tube 31 is not introduced into the external one 32 morethan one diameter of the external tube 32 and provided that the internaltube 31 does not project more than two diameters from the external tube32. The best results are obtained when the internal tube 31 projectsfrom the external one 32 a distance substantially the same as thediameter of the internal tube 31. This same criterion is valid if morethan two tubes are used, with the tube that is surrounded (inner tube)projecting beyond the tube that surrounds (outer tube) by a distancesubstantially the same as the diameter of the first tube.

The distance between the plane of the internal tube 31 (the one thatwill normally project more) and the plane of the orifice may varybetween zero and three outside diameters of the external tube 32,depending on the surface tensions between the liquids and with the gas,and on their viscosity values. Typically, the optimal distance is foundexperimentally for each particular configuration and each set of liquidsused.

The proposed atomizing system obviously requires fluids that are goingto be used in the resulting spray to have certain flow parameters.Accordingly, flows for this use must be:

-   -   Flows that are suitable so that the system falls within the        parametric window of stability. Multiplexing (i.e. several sets        of concentric tubes) may be used, if the flows required are        greater than those of an individual cell.    -   Flows that are suitable so that the mass relation of the fluids        falls within the specifications of each application. Of course,        a greater flow of gas may be supplied externally by any means in        specific applications, since this does not interfere with the        functioning of the atomizer.    -   If the flows are varied, the characteristic time of this        variation must be less than the hydrodynamic residence times of        liquid and gas in the microjet, and less than the inverse of the        first natural oscillation frequency of the drop formed at the        end of the injection needle.

Therefore, any means for continuous supply of gas (compressors, pressuredeposits, etc.) and of liquid (volumetric pumps, pressure bottles) maybe used. If multiplexing is desired, the flow of liquid must be ashomogeneous as possible between the various cells, which may requireimpulse through multiple capillary needles, porous media, or any othermedium capable of distributing a homogeneous flow among differentfeeding points.

Each atomizing device will consist of concentric tubes 31, 32 with adiameter ranging between 0.05 and 2 mm, preferably between 0.1 and 0.4mm, on which the drop from which the microjet emanates can be anchored,and a small orifice (between 0.001 and 2 mm in diameter, preferablybetween 0.1 and 0.25 mm), facing the drop and separated from the pointof feeding by a distance between 0.001 and 2 mm, preferably between 0.2and 0.5 mm. The orifice puts the suction gas that surrounds the drop, athigher pressure, in touch with the area in which the atomizing is to beattained, at lower pressure.

EMBODIMENT OF FIG. 3

The embodiments of FIGS. 1 and 2 are similar in a number of ways. Bothhave a feeding piece which is preferably in the form of a feeding needlewith a circular exit opening. Further, both have an exit port in thepressure chamber which is positioned directly in front of the flow pathof fluid out of the feeding source. Precisely maintaining the alignmentof the flow path of the feeding source with the exit port of thepressure chamber can present an engineering challenge particularly whenthe device includes a number of feeding needles. The embodiment of FIG.3 is designed to simplify the manner in which components are aligned.The embodiment of FIG. 3 uses a planar feeding piece (which by virtue ofthe withdrawal effect produced by the pressure difference across a smallopening through which fluid is passed) to obtain multiple microjetswhich are expelled through multiple exit ports of a pressure chamberthereby obtaining multiple aerosol streams. Although a single planarfeeding member as shown in FIG. 3 it, of course, is possible to producea device with a plurality of planar feeding members where each planarfeeding member feeds fluid to a linear array of outlet orifices in thesurrounding pressure chamber. In addition, the feeding member need notbe strictly planar, and may be a curved feeding device comprised of twosurfaces that maintain approximately the same spatial distance betweenthe two pieces of the feeding source. Such curved devices may have anylevel of curvature, e.g. circular, semicircular, elliptical,hemi-elliptical etc.

The components of the embodiment of FIG. 3 are as follows:

41. Feeding piece.

42. End of the feeding piece used to insert the fluid to be atomized.

43. Pressure chamber.

44. Orifice used as gas inlet.

45. End of the feeding needle used to evacuate the liquid to beatomized.

46. Orifices through which withdrawal takes place.

47. Atomizate (spray) or aerosol.

48. first fluid containing material to be atomized.

49. second fluid for creation of microjet.

50. wall of the propulsion chamber facing the edge of the feeding piece.

51. channels for guidance of fluid through feeding piece.

d_(j)=diameter of the microjet formed; ρ_(A)=liquid density of firstfluid (48); ρ_(B)=liquid density of second fluid (49); v_(A)=velocity ofthe first liquid (48); v_(B)=velocity of the second liquid (49); e=axiallength of the orifice through which withdrawal takes place; H=distancefrom the feeding needle to the microjet outlet; P₀=pressure inside thechamber;

Δp_(g)=change in pressure of the gas; P_(α)=atmospheric pressure;Q=volumetric flow rate

The proposed dispersing device consists of a feeding piece 41 whichcreates a planar feeding channel through which a where a first fluid 48flows. The flow is preferably directed through one or more channels ofuniform bores that are constructed on the planar surface of the feedingpiece 41. A pressure chamber 43 that holds the propelling flow of asecond liquid 49, houses the feeding piece 41 and is under a pressureabove maintained outside the chamber wall 50. One or more orifices,openings or slots (outlets) 46 made in the wall 52 of the propulsionchamber face the edge of the feeding piece. Preferably, each bore orchannel of the feeding piece 41 has its flow path substantially alignedwith an outlet 46.

Formation of the microjet and its acceleration are based on the abruptpressure drop resulting from the steep acceleration undergone by thesecond fluid 49 on passing through the orifice 46, similarly to theprocedure described above for embodiments of FIGS. 1 and 2 when thesecond fluid 49 is a gas.

When the second fluid 49 is a gas and the first fluid 48 is a liquid,the microthread formed is quite long and the liquid velocity is muchsmaller than the gas velocity. In fact, the low viscosity of the gasallows the liquid to flow at a much lower velocity; as a result, themicrojet is actually produced and accelerated by stress forces normal tothe liquid surface, i.e. pressure forces. Hence, one effectiveapproximation to the phenomenon is to assume that the pressuredifference established will result in the same kinetic energy per unitvolume for both fluids (liquid and gas), provided gas compressibilityeffects are neglected. The diameter d_(j) of the microjet formed from aliquid density ρ₁ that passes at a volumetric flow-rate Q through anorifice across which a pressure difference ΔP_(g) exists will be givenby

d j - ( 8 ⁢ ⁢ ρ l π 2 ⁢ Δ ⁢ ⁢ P g ) 1 / 4 ⁢ Q

See Gañàn-Calvo, Physical Review Letters, 80:285-288 (1998).

The relation between the diameter of the microjet, d_(j), and that ofthe resulting drops, _, depends on the ratio between viscous forces andsurface tension forces on the liquid on the one hand, and betweendynamic forces and surface tension forces on the gas on the other (i.e.on the Ohnesorge and Weber numbers, respectively) (Hinds (AerosolTechnology, John & Sons, 1982), Lefevre (Atomization and Sprays,Hemisphere Pub. Corp., 1989) and Bayvel & Orzechowski (LiquidAtomization, Taylor & Francis, 1993)). At moderate to low gas velocitiesand low viscosities the relation is roughly identical with that forcapillarity instability developed by Rayleigh:_(—)=1.89d_(j)

Because the liquid microjet is very long, at high liquid flow-rates thetheoretical rupture point lies in the turbulent zone created by the gasjet, so turbulent fluctuations in the gas destabilize or rupture theliquid microjet in a more or less uneven manner. As a result, thebenefits of drop size uniformity are lost.

On the other hand, when the second fluid 49 is a liquid and the firstfluid 48 is a gas, the facts that the liquid is much more viscous andthat the gas is much less dense virtually equalize the fluid and gasvelocities. The gas microthread formed is much shorter; however, becauseits rupture zone is almost invariably located in a laminar flowingstream, dispersion in the size of the microbubbles formed is almostalways small. At a volumetric gas flow-rate Q_(g) and a liquidoverpressure ΔP₁, the diameter of the gas microjet is given by

$d_{j} - {\left( \frac{8\rho_{l}}{\pi^{2}\Delta\; P_{l}} \right)^{1/4}Q_{g}^{1/2}}$

The low liquid velocity and the absence of relative velocities betweenthe liquid and gas lead to the Rayleigh relation between the diametersof the microthread and those of the bubbles (i.e. d=1.89d_(j)).

If both fluids 48, 49 are liquid and scarcely viscous, then theirrelative velocities will be given by

v A v B = ( ρ B ρ A )

The diameter of a microjet of the first liquid at a volumetric flow-rateof A Q_(A) and

$d_{j} - {\left( \frac{8\rho_{A}}{\pi^{2}\Delta\; P_{B}} \right)^{1/4}Q_{A}^{1/2}}$an overpressure of BΔP_(B) will be given by

At viscosities such that the velocities of both fluids 48, 49 willrapidly equilibrate in the microjet, the diameter of the microjet of thefirst liquid will be given by

$d_{j} - {\left( \frac{8\rho_{B}}{\pi^{2}\Delta\; P_{B}} \right)^{1/4}Q_{A}^{1/2}}$

The proposed atomization system obviously requires delivery of thefluids 48, 49 to be used in the dispersion process at appropriateflow-rates. Thus:

(1) Both flow-rates should be adjusted for the system so that they liewithin the stable parameter window.

(2) The mass ratio between the flows should be compatible with thespecifications of each application. Obviously, the gas flow-rate can beincreased by using an external means in special applications (e.g.burning, drug inhalation) since this need not interfere with theatomizer operation.

(3) If the flow-rates are altered, the characteristic time for thevariation should be shorter than the hydrodynamic residence times forthe liquid and gas in the microjet, and smaller than the reciprocal ofthe first natural oscillation frequency of the drop formed at the end ofthe feeding piece.

(4) Therefore, the gas and liquid can be dispensed by any type ofcontinuous delivery system (e.g. a compressor or a pressurized tank theformer and a volumetric pump or a pressurized bottle the latter).

(5) The atomizer can be made from a variety of materials (metal,plastic, ceramics, glass).

DRUG DELIVERY DEVICES

The various embodiments of components for creating aerosols describedabove can be used in devices for the delivery of an aerosol to apatient. The device is preferably a hand-held, self-contained devicewhich a patient can easily carry about and use for the administration ofdrugs.

In one embodiment the pressurized canister of a device as disclosed inU.S. Pat. Nos. 5,364,838 or 5,404,871 is replaced with an aerosolgenerating device of the type shown here in any of FIGS. 1, 2 or 3. Thedevice preferably includes a means for measuring a patient's respiratoryflow rate and respiratory volume. The aerosol may be generated when thepatient manually actuates release of aerosol. However, the devicepreferably operates as the device in the U.S. Pat. No. 5,364,838 patentand is actuated automatically in response to a measured inspiratory flowrate and inspiratory volume. This makes it possible to repeatedlydeliver aerosol to a patient at the same point in the respiratory cyclethereby improving repeatability of dosing.

In another embodiment the disposable containers and porous membranes ofthe device disclosed in U.S. Pat. No. 5,544,646 is replaced with anaerosol generating device of the type shown here in FIGS. 1, 2 or 3. Insuch an embodiment the aerosol generating device is likely to include aplurality of sources of liquid drug formulation (e.g. a plurality offeeding needles) so that a sufficient amount of formulation can beaerosolized in a sufficiently short period of time that the patient can,in a single inhalation, inhale all of the aerosol needed for a dose ofthe drug being delivered. An aerosol generating device of the type shownin FIG. 3 is generally easier to manufacture and as such is preferred inthis embodiment.

In the two embodiments described above it is preferable to create anaerosol with particles having a particle diameter in the range of about1 to about 3 microns. This allows the particles to penetrate into thesmallest channels of the lungs. However, such particles are not so smallthat they will not settle on the surface of these channels wheninhaled—failure to settle would cause the particles to be exhaled aswith particles of smoke.

In yet another embodiment the source of aerosol of a device as disclosedin U.S. Pat. No. 4,484,577 is replaced with an aerosol generating deviceof the type described here and shown in FIGS. 1, 2 or 3. The device ofU.S. Pat. No. 4,484,577 includes a large container into which theaerosol is dispersed. The aerosol then “hangs” in the air in the largecontainer until it is inhaled by the patient. Using an aerosolgeneration device of the invention the aerosol particles can be madevery small, e.g. about 1 micron or less. Particles of this size will besuspended in air in the container until inhaled. The patient can thenbreath normally in and out of the container until all or substantiallyall particles in the container have deposited themselves on the channelsof the lung.

EXAMPLES

The following examples are put forth so as to provide those of ordinaryskill in the art with a complete disclosure and description of how tomake and use the present invention, and are not intended to limit thescope of what the inventors regard as their invention nor are theyintended to represent that the experiments below are all or the onlyexperiments performed. Efforts have been made to ensure accuracy withrespect to numbers used (e.g. amounts, temperature, etc.) but someexperimental errors and deviations should be accounted for. Unlessindicated otherwise, parts are parts by weight, molecular weight isweight average molecular weight, temperature is in degrees Centigrade,and pressure is at or near atmospheric.

The properties of sixteen different liquids are provided in Table 1

TABLE 1 Liquids used and some of their physical properties at 24.5° C.(ρ: kg/m³, μ: cpoise, γ: N/m). Also given, the symbols used in theplots. Liquid ρ μ γ Symbol Heptane 684 0.38 0.021

Tap Water 1000 1.00 0.056 ⋄ Water + glycerol 90/10 v/v 1026 1.39 0.069 −Water + glycerol 80/20 v/v 1052 1.98 0.068 ∇ Isopropyl alcohol 755.52.18 0.021 X Water + glycerol 70/30 v/v 1078 2.76 0.067 0 Water +glycerol 60/40 v/v 1104 4.37 0.067 • Water + glycerol 50/50 v/v 10306.17 0.066

1-Octanol 827 7.47 0.024 ⋄ Water + glycerol 40/60 v/v 1156 12.3 0.065 −Water + glycerol 35/65 v/v 1167 15.9 0.064 ∇ Water + glycerol 30/70 v/v1182 24.3 0.064 X Water + glycerol 25/75 v/v 1195 38.7 0.063 + Propyleneglycol 1026 41.8 0.036 •

The liquids of Table 1 were forced through a feeding needle of the typeshown in FIG. 1. The end 5 of the feeding needle had an internal radiusR_(o). The exit orifice 6 had a diameter D and the wall of the pressurechamber 3 had a thickness of L. Three different devices were testedhaving the following dimensions: (D=0.15, 0.2, and 0.3 mm; L=0.1, 0.2and 0.35 mm; R_(o)+0.2, 0.4, and 0.6 mm, respectively), and severaldistances H from the tube mouth to the orifice ranging from H=0.5 mm toH=1.5 mm have been used. The jet diameter was measured at the hole exitand was plotted as a function of the pressure difference ΔP_(g) and flowrate Q respectively. Although this technique allows for jet diameterseven below one micron, larger flow rates and diameters have been used inthis study to diminish the measuring errors.

In order to collapse all of the data, we define a reference flow rateQ_(o) and diameter d_(o) based on the minimal values, from expressions(3) and (5), that can be attained in stable regime for a given ΔP_(g):

Q o = ( γ 4 ρ l ⁢ Δ ⁢ ⁢ P g 3 ) , d o = γ Δ ⁢ ⁢ P g ( 6 )

These definitions provide the advantage of a nondimensional expressionfor (5), asd _(j) /d _(o)=(8/π²)^(1/4)(Q/Q _(o))^(1/2),  (7)

which allows for a check for the validity of neglecting the surfacetension term in (4) (i.e., Q/Q_(o) should be large).

Notice that if the measured d_(j) follows expression (5), the surfacetension cancels out in (7). Also notice that d_(j)/d_(o)≅We/2.

350 measured values of d_(j)/d_(o) versus Q/Q_(o) are plotted in FIG. 5.A continuous line represents the theoretical prediction (7), independentof liquid viscosity and surface tension. The use of different hole andtube diameters as well as tube-hole distances does not have anyappreciable influence on d_(j). The collapse of the experimental dataand the agreement with the simple theoretical model is excellent.Finally, the experimental values of Q are at least four times large thanQ_(o) (being in most cases several hundreds times larger), whichjustifies the neglect of the surface tension term in Eq. (4).

While the present invention has been described with reference to thespecific embodiments thereof, it should be understood by those skilledin the art that various changes may be made and equivalents may besubstituted without departing from the true spirit and scope of theinvention. In addition, many modifications may be made to adapt aparticular situation, material, composition of matter, process, processstep or steps, to the objective, spirit and scope of the presentinvention. All such modifications are intended to be within the scope ofthe claims appended hereto.

1. A method comprising the steps of: feeding a pharmaceutically activedrug through a cylindrical channel of a feeding source in a manner whichcauses the liquid to be expelled from an exit opening as a liquid streamwherein the exit opening has a diameter in a range of from about 0.002mm to about 2 mm; forcing a gas through a pressure chamber in a mannerwhich causes the gas to exit the pressure chamber from an exit orificehaving a diameter in a range of about 0.002 mm to about 2 mm, the gasexiting downstream of a flow path of the liquid stream expelled from theexit opening of the feeding source; wherein the exit opening of thefeeding source is separated by a distance of from about 0.002 mm toabout 2 mm from the exit opening of the feeding source; wherein theliquid is forced through the channel at a rate in a range of about 0.01nl/sec to about 100 microliters/sec.
 2. The method of claim 1, furthercomprising: allowing the liquid stream to form particles.
 3. The methodof claim 2, wherein the particles have an aerodynamic diameter in arange of from 0.1 micron to about 10 microns.
 4. The method of claim 2,wherein the particles have an aerodynamic diameter in a range of from 1to 5 microns.
 5. The method of claim 2, wherein the particles have thesame diameter with a relative standard deviation of 10% to 30%.
 6. Themethod of claim 2, wherein the particles have the same diameter with arelative standard deviation of 3% to 20%.
 7. The method of claim 1,wherein the pharmaceutically active drug is a solution.
 8. The method ofclaim 1, wherein the pharmaceutically active drug is a suspension. 9.The method of claim 1, wherein the liquid has a viscosity in a range offrom about 10⁻⁴ to about 1 kg/m/sec.
 10. The method of claim 1, whereinthe gas is a non-toxic gas.
 11. The method of claim 10, wherein thenon-toxic gas is CO₂.
 12. The method of claim 10, wherein the liquidformulation fed through the channel has a volume in the range of 5microliters to 10,000 microliters.
 13. The method of claim 12, whereingas is forced into an area around the feeding source outlet at apressure below 50,000 mBar above atmospheric pressure.
 14. The method ofclaim 1, wherein the gas is air.
 15. The method of claim 1, wherein thegas is forced through the opening of the pressure chamber at a rate inthe range of from about 50 m/sec to about 2000 m/sec.
 16. The method ofclaim 1, wherein the exit opening has a diameter in the range of fromabout 0.01 mm to about 0.4 mm.
 17. The method of claim 1, wherein theliquid formulation fed through the channel has a volume in the range of5 microliters to 10,000 microliters.
 18. The method of claim 1, whereingas is forced into an area around the feeding source outlet at apressure above 10 mBar above atmospheric pressure.
 19. The method ofclaim 1, wherein the liquid and gas are each fed at a rate relative toeach other so as to form aerosolized particles having a size in therange of about 0.1 micron to about 10 microns.
 20. The method of claim1, wherein the liquid has a viscosity in a range of from about 0.3×10⁻³to about 5×10⁻² kg/m/sec; wherein the liquid is forced through thecylindrical channel at a rate in a range of about 0.01 nl/sec to about100 μl/sec and further wherein the gas is forced through the exitorifice of the pressure chamber at a rate in the range of from about 50m/sec to about 2000 m/sec.
 21. The method of claim 1, wherein the liquidis forced through the cylindrical channel at a rate in a range of about1 nl/sec to about 10 μl/sec and further wherein the gas is forcedthrough the exit orifice of the pressure chamber at a range in a rangeof from about 100 to 500 m/sec.
 22. The method of claim 1, wherein theexit opening has a diameter in the range of from about 0.01 mm to about0.4 mm, and and wherein the exit opening of the feeding source isseparated by a distance of from about 0.01 to about 2 mm from the exitopening in the pressure chamber.
 23. The method of claim 1, wherein theopening in the pressure chamber has a diameter in the range of about0.005 mm to about 0.25 mm, and wherein the exit opening of the feedingsource is separated by a distance of from about 0.002 to about 2 mm fromthe feeding point.
 24. The method of claim 1, wherein gas is forced intoan area around the feeding source outlet at a pressure in the range of10 to 50,000 mBar above atmospheric pressure and further wherein theliquid has a viscosity in the range of from 10⁻⁴ to 1 kg/m/sec.
 25. Themethod of claim 24, wherein gas is forced into an area around thefeeding source outlet at a pressure in the range of 100-2000 mBar aboveatmospheric pressure.
 26. The method of claim 1, wherein gas from thepressure chamber surrounds liquid exiting the feeding source exitopening which liquid is concentrically focused by the gas flowing out ofthe exit orifice and further wherein the aerosolized particles formedare uniform in size to the extent of having a relative size standarddeviation of 3 to 30%.